Positron Emission Tomography (PET) is a proven and powerful research and clinical tool that uses molecular biomarkers for imaging the function and metabolism of the human body and brain. These biomarkers can be tailored to a high degree of specificity to highlight abnormal metabolism and uptake and are often used to diagnose diseases such as brain tumor, strokes, neuron-damaging disease (dementia), and other neurological disorders (addiction, schizophrenia, etc.). However, it suffers a number of critical limitations that restrict its full research and clinical potential.
The inability to rapidly diagnosis non-penetrating Traumatic Brain Injury (TBI) especially in its mild form (mTBI) has very significant implications for military operations, and for the health and health care costs of TBI victims. Detection of TBI is difficult as symptoms can be slow to develop and are difficult to recognize. Premature return to duty of an effected individual could seriously impact on his/her performance, and delayed medical treatment exacerbates an individual's acute and chronic medical condition.
Present human brain PET scanners are large, heavy, and fragile with ring geometries and hence are of low sensitivity. They have low and non-uniform spatial resolution and very poor temporal resolution. None of them allows imaging of a freely moving subject‡ [31, 32], only one of them allows imaging in a sitting position (PET-Hat)[33]. Other all require subjects to be lying on a scanner bed. This restriction limits the use of PET brain mapping in studies of, for example, bipolar disorder [34-46], psychosis, schizophrenia, autism, dementia, glutamate levels, obesity [47-50], gambling [51], video game addiction [52], and chronic fatigue syndrome [53-55]. Additionally, all conventional PET devices are sensitive to environmental factors (temperature gradients, vibrations, etc.).
U.S. Pat. No. 7,884,331, expressly incorporated herein by reference, provides a brain imager which includes a compact ring-like static PET imager mounted in a helmet-like structure, as shown in FIG. 8. When attached to a patient's head, the helmet-like brain imager maintains the relative head-to-imager geometry fixed through the whole imaging procedure. The brain imaging helmet contains radiation sensors and minimal front-end electronics. A flexible mechanical suspension/harness system supports the weight of the helmet thereby allowing for patient to have limited movements of the head during imaging scans. The compact ring-like PET imager enables very high resolution imaging of neurological brain functions, cancer, and effects of trauma using a rather simple mobile scanner with limited space needs for use and storage. This system provides a ring of detectors, in a single layer, in a non-wearable helmet (suspended by a cable), and is not conformable or adapted to form arbitrary 3D shaped arrays, and does not provide depth of interaction information.
The Stanford University Molecular Imaging Instrumentation Laboratory developed a scintillation light detection system for PET. The system is constructed with basic block detector modules, for clinical or pre-clinical (small animal) PET systems and breast-dedicated PET system. The module is in the form of many detector layers. Each layer comprises two adjacently situated planar arrays of 1×1×1 mm3 lutetium-ytrium-oxyorthosilicate (LYSO) scintillation crystals coupled to two specially designed, planar and extremely thin (200 micron) position-sensitive avalanche photodiodes (PSAPD), each with a 8×8 mm2 sensitive area. Alternating layers of planar crystal arrays and flat PSAPD detectors, coupled together, are configured “edge-on” with respect to incoming photons. The PSAPD has both high light sensitivity and an intrinsic resolution that is finer than 1 mm, the selected width of the LYSO crystals. The PSAPD generates large electronic signals and allows the precise positioning of the light flashes resulting from the absorption of a 511 keV in any crystal; the resolution of each LSO-PSAPD compound layer in any direction is determined primarily by the 1-mm individual crystal dimension. The PSAPDs replace the photomultiplier tubes (PMT) used in standard PET designs. The PSAPD reads the crystals from their relatively large side faces, rather than from their tiny end faces for better scintillation light collection efficiency and directly-measured photon interaction depth. This approach allows high efficiency light extraction from the long and thin scintillation crystals; with light collection efficiency >95%. Light collection per event is largely independent of crystal length or surface treatment, and of location within a crystal of the point of light creation. Coincidence detection efficiency (i.e., fraction of emitted coincident 511 keV photon counts collected from a given probe concentration) is increased by bringing the detectors closer to the subject and by using thicker scintillation crystal material (2 cm effective thickness proposed vs. 1 cm standard for small animal imaging). Spatial resolution is enhanced by using finer scintillation crystals (1 mm vs. 1.5-6.0 mm standard). Uniformity of spatial resolution improves by direct measurement of the photon interaction depth within the crystals using the side-coupled PSAPDs; this depth measurement reduces position-dependent parallax positioning errors (hence loss of resolution) due to photon penetration into crystals. Standard PET system detectors are incapable of this photon interaction depth resolution. See, miil.stanford.edu/research/scintillationlightdetection.html; miil.stanford.edu/publications/files/39_PUB.pdf; 29_PUB.pdf; 24_PUB.pdf; 14_PUB.pdf; 142_PUB.pdf; 183_PUB.pdf; 147_PUB.pdf; 52_PUB.pdf; 181_PUB.pdf; each of which is expressly incorporated herein by reference. U.S. Pat. No. 8,373,132, expressly incorporated herein by reference, relates to a radiation detector with a stack of scintillator elements and photodiode arrays. The detector comprises a stack of the scintillator elements and photodiode arrays (PDA). The PDAs extend with electrical leads into a rigid body filling a border volume lateral of the scintillator elements, wherein said leads end in a contact surface of the border volume. Moreover, a redistribution layer is disposed on the contact surface, wherein electrical lines of the redistribution layer contact the leads of the PDAs.
US 20110240864, expressly incorporated herein by reference, discloses an autonomous detector module as a building block for scalable PET and SPECT systems. When detecting scintillation events in a nuclear imaging system, time-stamping and energy-gating processing is incorporated into autonomous detection modules (ADM) to reduce downstream processing. Each ADM is removably coupled to a detector fixture, and comprises a scintillation crystal array and associated light detect or (s), such as a silicon photomultiplier or the like. The light detector(s) is coupled to a processing module in or on the ADM, which performs the energy gating and time-stamping.
SynchroPET (Long Island, N.Y.) is seeking to commercialize a portable, small-scale brain-imaging device invented at the U.S. Department of Energy's Brookhaven National Laboratory. (Justin Eure, Groundbreaking Portable PET Scanner Moves Closer to Market, Medical Applications, PHYS.ORG, Jan. 27, 2012 phys.org/news/2012-01-groundbreaking-portable-pet-scanner-closer.html.). The mini PET scanner is intended to be integrated with an MRI. It also has been worn like a collar by fully conscious, active rats. (Miniature “Wearable” PET Scanner Ready for Use, BROOKHAVEN NATIONAL LABORATORY, Mar. 13, 2011 www.bnl.gov/newsroom/news.php?a=11235.) This miniature PET scanner seeks to integrate all the electronics for each detector in the ring on a single, specialized chip, using an avalanche photodiode. (24 Katherine Bourzac, Wearable Scanner Opens New Frontier in Neuroscience, MIT TECHNOLOGY REVIEW, Mar. 17, 2011, m.technologyreview.com/biomedicine/35127/.)
U.S. Pat. No. 7,286,867, expressly incorporated herein by reference, directed to the SynchroPET device provides a combined PET/MRI scanner generally includes a magnet for producing a magnetic field suitable for magnetic resonance imaging, a radiofrequency (RF) coil disposed within the magnetic field produced by the magnet and a ring tomograph disposed within the magnetic field produced by the magnet. The ring tomograph includes a scintillator layer for outputting at least one photon in response to an annihilation event, a detection array coupled to the scintillator layer for detecting the at least one photon outputted by the scintillator layer and for outputting a detection signal in response to the detected photon and a front-end electronic array coupled to the detection array for receiving the detection signal, wherein the front-end array has a preamplifier and a shaper network for conditioning the detection signal. One embodiment of this technology comprises a 4×8 array of LSO crystals of about 2 mm×2 mm square area and 10 mm depth. Another embodiment includes a 5 mm thick LSO scintillator array for better spatial resolution. However, the spatial resolution is still only 2.5 mm to 1.9 mm, respectively. Although this patent describes that freedom of movement is available to the animal, its appears to require a suspension system.
Another company working in the portable PET scanner space is the Brain Biosciences. (Wayne Forrest, Start-up Brain Biosciences Advances Compact PET Scanner, AUNTMINNIE.COM, Jan. 23, 2014 www.auntminnie.com/index.aspx?sec=ser&sub=def&pag=dis&ItemID=106270.) which developed the CerePET, a portable PET scanner for neuroimaging. The prototype is approximately 50 lbs. and can be carried from room to room. The device features crystals made from lutetium yttrium orthosilicate (LYSO) technology, a bore diameter of 25 cm, and 20-cm axial field-of-view. However, the device is not wearable. The scanner has a head support system that fits on a standard examination bed. Preclinical work with phantoms showed that the system is capable of spatial resolution of 2 mm to 3 mm across the field-of-view, energy resolution of less than 13% for all detector blocks, and more than 15% image uniformity. Quantitative accuracy was also better than 10% after calculating attenuation correction.
WO 2012/034178, expressly incorporated herein by reference, provides a radiation detector is disclosed that comprises a scintillator that emits electromagnetic radiation in response to excitation by radiation of interest, and a photodetector with a semiconductor active layer adapted to interact with the electromagnetic radiation and a substrate. Either the scintillator constitutes the substrate or the substrate is located such that the substrate is between said active layer and the scintillator, and the substrate is at least partially transparent to the electromagnetic radiation.
U.S. Pat. No. 8,071,949, expressly incorporated herein by reference, provides a compact, mobile, dedicated SPECT brain imager that can be easily moved to the patient to provide in-situ imaging, especially when the patient cannot be moved to the Nuclear Medicine imaging center. As a result of the widespread availability of single photon labeled biomarkers, the SPECT brain imager can be used in many locations, including remote locations away from medical centers. The SPECT imager improves the detection of gamma emission from the patient's head and neck area with a large field of view. Two identical lightweight gamma imaging detector heads are mounted to a rotating gantry and precisely mechanically co-registered to each other at 180 degrees. A unique imaging algorithm combines the co-registered images from the detector heads and provides several SPECT tomographic reconstructions of the imaged object thereby improving the diagnostic quality especially in the case of imaging requiring higher spatial resolution and sensitivity at the same time.
U.S. Pat. No. 7,554,089, expressly incorporated herein by reference, provides a method for localizing optical emission is disclosed. The method involves identifying a first readout channel of a first pixellated photodetector array based on an impact of a first photon on the first pixellated photodetector array. The first photon is emitted by a scintillator unit of a scintillator array and the first readout channel corresponds to a column of one or more pixels of the first pixellated photodetector array. The method also involves identifying a second readout channel of a second pixellated photodetector array based on an impact of a second photon on the second pixellated photodetector array. The second photon is emitted by the scintillator unit and the second readout channel corresponds to a row of one or more pixels of the second pixellated photodetector array. The method further involves identifying the scintillator unit based on the first readout channel and the second readout channel.
US 2013/0153774, expressly incorporated herein by reference, provides apixellated scintillator readout arrangement is presented, the arrangement comprising a plurality of scintillator pixels arranged in a scintillator array, and a plurality of photodetectors arranged to receive light from, or address, the scintillator pixels. The photodetectors may be arranged on both a first side and a second side of the scintillator array. Each photodetector may be arranged to leave a gap adjacent to the scintillator pixel which is addressed by that photodetector. Non-photosensitive elements such as tracking and bondpads may be arranged in at least some of the gaps. Electronic components such as electronic amplifiers may be arranged in at least some of the gaps. The photodetectors may be arranged in linear arrays addressing alternate lines of scintillator pixels on either side of the scintillator array. Each photodetector may be arranged to address a single pixel (as illustrated) or more than one pixel.
U.S. Pat. No. 7,956,331, expressly incorporated herein by reference, provides scintillation detectors capable of detecting the position or depth of gamma photon interactions occurring within a scintillator, thereby improving the resolution of ring based positron emission tomography (PET) imaging systems. In one embodiment, the technology is directed to a scintillation detector that comprises at least one pair of side-by-side conjunct scintillation crystal bars having a shared interface between, and a solid-state semiconductor photodetector optically coupled to each output window of each individual scintillation crystal bar. The solid-state semiconductor photodetector includes an array of discrete sensitive areas disposed across a top surface of a common substrate, wherein each sensitive area contains an array of discrete micro-pixelated avalanche photodiodes, and wherein the output window of each scintillation crystal bar is optically coupled to each respective sensitive area in a one-on-one relationship. This patent describes scintillation detectors capable of detecting the position or depth of gamma photon interactions occurring within a scintillator. At least one pair of conjunct scintillation crystal bars may be provided, wherein each pair of conjunct scintillation crystal bars is composed of two individual optically coupled scintillation crystal bars positioned in a side-by-side relationship. The scintillation detector also includes a solid-state semiconductor photodetector.
WO 2012152587, expressly incorporated herein by reference, provides a Gamma Detector which comprises a scintillation crystal block and a set of Geiger-mode Avalanche Photodiode (G-APD) sensor elements optically coupled to at least a first surface of the scintillation crystal block. The G-APD sensor elements are arranged in at least one elongate strip of G-APD sensor elements, said G-APD strip coupled to a readout circuit.
The emission of the scintillator must be detected with an energy resolution sufficient to distinguish the 511 keV primary events from scattering events and with timing resolution to enable coincidence windowing. For a PET scanner, the requirements also include size and scalability, operation in high magnetic fields for multimodality imaging, and cost; these pose substantial challenges to the system design[63]. Photomultiplier have traditionally been used as they provide high gain (106) with fast response and low noise, can be fabricated as arrays for segmented detection, and are relatively low cost. However, the physical size limits use of PMT based detectors in applications where detector volume and mass must be minimized, and they cannot operate in high magnetic fields for MRI.
Solid state detectors address many of the limitations of PMTs and numerous configurations have been reported. One approach is direct conversion and detection, as recently demonstrated in CdTe Schottky barrier detectors in 4×4×2 mm crystals [64]. However, all direct conversion techniques require high voltage and are difficult to scale to large arrays. Consequently, the majority of solid-state detectors utilize a scintillator crystal coupled to the thin film photon detection sensors (TFPDS), primarily Si PIN diodes, avalanche photodiodes (APD), and Geiger-mode avalanche photodiodes (G-APD). State of the art Si-based solid state photon detectors were reviewed in 2009 by Renker and Lorentz [65]. For PET applications, G-APD detectors show the greatest promise as they provide the energy resolution, speed, and integration level required for normal and time-of-flight detection. Arrays of G-APDs can be fabricated over relatively large areas (1×1 cm), as demonstrated by Radiation Monitoring Devices [rmdinc.com/solid-state-photo-multipliers-for-pet/]. Further, coupled with a LaBr3:Ce scintillator, timing resolution to 100 ps has been demonstrated providing potential for time-of-flight enhancements to image reconstruction [66].
Table 2 provides incomplete list of scintillators suitable for PET scanners [58-62]. NaI(T1) has been used in PET designs in the past in view of its brighter response, low cost and good energy resolution. However, its slow response and low-gamma ray stopping efficiency limit its performance in PET. Europium doped strontium iodide (SrI2:Eu) has very high light output and excellent energy resolution for gamma detection, but rather slow response due to high decay time constant. LSO and LYSO are used in PET instrumentation due to their high gamma ray stopping power. Still, energy resolution of LSO is variable and is limited by its non-proportionality (self-absorption). Furthermore, LSO is optically anisotropic, and it cannot be made fully transparent. Lutetium-yttrium orthosilicate (LYSO) is a variant of LSO in which some of the lutetium is replaced by yttrium atoms and is currently the scintillator of choice for PET application. The yttrium component makes LYSO easier and cheaper to grow (low temperature crystal growth), and the light loss from self-absorption is lower in LYSO. Due to the presence of Lu-176 isotope, LSO/LYSO crystals are themselves radioactive and produce scintillation light. Lu-176 is a β-emitter primarily decaying to an excited state of Hf-176, which then emits gamma photons with energies of 307 keV, 202 keV, and 88 keV [61]. Lanthanum bromide LaBr3:Ce crystals have the advantages of very high light output (˜3× of LYSO), fast decay time affording excellent time resolution, excellent energy resolution (<3% at 662 keV), thus providing improved ability for rejection of scattered events and random coincidences, and low melting point (783° C., compared to 2047° C. for LYSO) offering easier crystal growth at lower cost.
In a PET system the ultimate goal is to produce a high fidelity image of the radionuclide distribution within the imaged object. There are three fundamental mechanisms that contribute to the quality of PET images: system sensitivity, noise, and spatio-temporal localization. [59]. The first two quantities are well accounted for in the data via the noise equivalent counts (NEC) [170]. All other things being equal the system with the better NEC will produce better images. Spatio-temporal localization is the system's basic ability to identify the spatio-temporal coordinates of the detected gamma photons (counts) and the necessary sampling to project them into the spatial domain (i.e. image reconstruction). It should be emphasized that this act of projecting the detected gamma photons into the spatial domain, the image reconstruction, is critical to achieving a high quality images even if it is not necessarily a part of the main system design process. We also note that improved localization (e.g. time-of-flight information) improves the effective NEC [171].
Image quality is directly related to the effective NEC collected during an acquisition, which includes improved noise rejection from time-of-flight (TOF) information [171]. It is defined by the relation,
                              NEC          eff                =                              (                          D                              min                ⁡                                  (                                                            c                      ⁢                                                                                          ⁢                      Δ                      ⁢                                                                                          ⁢                                              t                        /                        2                                                              ,                    D                                    )                                                      )                    ⁢                      (                                          T                +                S                +                                                      (                                          D                      /                                              D                        FOV                                                              )                                    ⁢                  R                                                            T                +                S                +                                                                            (                                              D                        /                                                  D                          FOV                                                                    )                                        2                                    ⁢                  R                                                      )                    ⁢          NEC                                        =                              (                          D                              min                ⁡                                  (                                                            c                      ⁢                                                                                          ⁢                      Δ                      ⁢                                                                                          ⁢                                              t                        /                        2                                                              ,                    D                                    )                                                      )                    ⁢                      (                                          T                +                S                +                                                      (                                          D                      /                                              D                        FOV                                                              )                                    ⁢                  R                                                            T                +                S                +                                                                            (                                              D                        /                                                  D                          FOV                                                                    )                                        2                                    ⁢                  R                                                      )                    ⁢                                    T              2                                      T              +              S              +              R                                          
where T, R, and S are true, random, and scatter counts respectively, and cΔt/2 represents the TOF photon localization, D is the mean head diameter, and DFOV is the diameter of the field-of-view (FOV). Clearly improved TOF for a given source geometry always improves the NEC.
To maximize the NEC for a particular activity distribution within the FOV a number of issues must be considered. They include geometric and attenuation sensitivity, detector dead time, source attenuation, sensitivity, and additive noise from random and scatter counts (and cascade with some isotopes). The geometric and attenuation sensitivity are defined by the device geometry, and the scintillator attenuation and thickness. On the other hand, dead time is dependent on the incident photon flux and the detector electronics. In general, the source's attenuation loss is patient dependent and cannot easily be included in the design consideration, however in the case of brain imaging, patient head size variability is smaller than variability of body habitus. Thus, its effects on sensitivity can be considered in the system's design.
The geometric sensitivity is defined by the effective solid angle of a point in the FOV and the detector scintillator. Because PET is based on coincidence detection, the geometric sensitivity is defined as 4π srd minus the solid angle subtended by the point to the non-detector area or opening and its back projection [172]. For systems with more than one opening care must be taken to avoid double counting any overlap between various solid angles and back projections. It follows that for spherical caps and cylindrical geometries the poorest geometric sensitivity in any particular slice is on the system's central axis. Detector attenuation sensitivity is defined by the system geometry and for a system with identical detectors it is can be well approximated by a constant.
The intrinsic resolution is largely a function of detection localization, and source uncertainty. The one aspect of detection localization is the system geometry, which is essentially determined by the solid angle coverage of the FOV and provides the available oversampling. It provides the overall detector framework and the subdivision of the detector elements into groups or blocks. Within this framework the light sharing and inter-detector Compton scatter define the ability of the detector to localize a particular detection event. These phenomena provide the detector's contribution to the overall system PSF [60]. In addition, depth of interaction (DOI) for off-center events can result in resolution loss for the system. The intrinsic source uncertainty is isotope dependent and arises from positron transport (decay location uncertainty, positron range) and residual momentum in the annihilation process (annihilation location uncertainty, non-collinearity). These sources of resolution uncertainty are independent and can be added in quadrature giving
            σ      total        =                            s          det          2                +                  σ                      pos            ⁢                          ,              A              Z                        ⁢            X                    2                +                  σ          θ          2                +                  s          ∥          2                      ,
where sdet
      s    det    ,      σ          pos      ⁢              ,        A        Z            ⁢      X        ,      σ    θ    ,          ⁢      and    ⁢                  ⁢          s      ∥      and s∥ are FWHM detector, positron range, non-collinearity and DOI uncertainties respectively [59]. We note that the positron range uncertainty is tissue and isotope dependent (density dependence provides a good approximation). The source location uncertainty can be found in the literature or estimated via Monte Carlo.
The detector's position uncertainty can be broken down further into detections via photoelectric and Compton interactions. The number of detected events for each type of interaction depends on their respective cross-section ratios for the particular interaction. In the case of the ADA system described here, where there is no light sharing; the uncertainty of detection via photoelectric effect (and correctly identified points of first interaction for Compton interactions) can be well approximated by the detector sampling distance because the range of the photoelectrons is very small, as compared to the size of scintillator crystals. The position uncertainty is limited by sampling and is sdet2=sPE2=(d/2)2, where d is the scintillator crystal width.
The effect of Compton scatter on resolution is more difficult to estimate and it cannot simply be added in quadrature to the sampling uncertainty. Nonetheless, it does contribute to the system point spread function. The uncertainty induced by Compton scatter is given by sc=√{square root over (xc2+d2/2)}, where the ½ term results from sampling and xc, is the mean-free-path of a 511 keV photon in the scintillator. The mean Compton path length can be derived and calculated via the integral
                    x        _            c        =                  ∫        0        π            ⁢                                    p            511                    ⁡                      (            θ            )                          ⁢                              sin            ⁢                                                  ⁢            θ                                              ρ              ⁡                              (                                  μ                  ρ                                )                                      E                          ⁢                                  ⁢        d        ⁢                                  ⁢        θ              ,
where p511(θ) is the probability of a particular scattering angle for incident 511 keV photons using Klein-Nishina cross sections, the sine term is a forward scattering correction, and xmfp=1/ρ(μ/ρ)E is the mean free path for a photon at a particular energy, which in this case is given by the Compton energy loss, E=−0.511/(2−cos θ).
In a multi-layered detector design, where individual detector elements cannot localize events, some DOI resolution loss will exist. The uncompensated loss for scintillators within blocks can be described by s∥=(xmfp/2)×(r√{square root over (R2−r2/R2)}), where r and R are the radial offset and system radius respectively. The factor of ½ is due to detector penetration likelihood at both ends of the line of response.
Depth of interaction correction can be accomplished by a number of means including position sensitive detectors and stacked detectors. In either case the resolution loss is mitigated by reducing the penetration depth effect. In the case of the stacked detector block the achievable FWHM resolution as a function of thickness results in the following relationship,
            σ      total        ⁡          (      t      )        =                    s        det        2            +              σ                  pos          ⁢                      ,            A            Z                    ⁢          X                2            +              σ        θ        2            +                        (                                                    min                ⁡                                  (                                      t                    ,                                                                  x                        _                                            mfp                                                        )                                            ⁢              r              ⁢                                                                    R                    2                                    -                                      r                    2                                                                                      2              ⁢                                                          ⁢                              R                2                                              )                2            
Note that the thickness must be less than the mean-free-path length of the photon in the scintillator.
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7,359,535; 7,365,333; 7,381,958; 7,381,959; 7,394,053; 7,405,405; 7,412,280; 7,447,345; 7,465,927; 7,489,799; 7,507,968; 7,519,412; 7,557,351; 7,576,329; 7,579,599; 7,605,373; 7,626,389; 7,638,771; 7,667,199; 7,667,457; 7,671,339; 7,684,589; 7,700,003; 7,705,314; 7,723,694; 7,737,404; 7,755,054; 7,756,310; 7,759,625; 7,778,787; 7,791,029; 7,800,070; 7,807,974; 7,818,047; 7,825,384; 7,847,552; 7,945,079; 7,953,265; 7,968,852; 7,983,735; 8,000,513; 8,003,948; 8,014,614; 8,017,902; 8,017,914; 8,068,896; 8,084,742; 8,094,908; 8,098,916; 8,110,805; 8,110,806; 8,144,962; 8,153,983; 8,155,415; 8,164,063; 8,193,815; 8,194,937; 8,204,172; 8,229,199; 8,258,480; 8,269,177; 8,274,054; 8,299,440; 8,304,736; 8,309,932; 8,315,353; 8,334,697; 8,340,377; 8,343,509; 8,350,220; 8,355,551; 8,357,486; 8,369,928; 8,384,037; 8,388,931; 8,395,127; 8,399,848; 8,405,035; 8,410,449; 8,431,904; 8,450,692; 8,450,693; 8,466,419; 8,467,848; 8,472,683; 8,472,688; 8,476,593; 8,476,594; 8,478,015; 8,481,947; 8,488,857; 8,497,484; 8,511,894; 8,525,116; 8,527,034; 8,530,846; 8,532,357; 8,547,100; 8,577,114; 8,598,534; 8,598,536; 8,604,440; 8,604,795; 8,605,988; 8,674,312; 8,698,087; 8,699,771; 8,716,664; 8,716,669; 8,723,521; 8,735,834; 8,735,835; 8,755,586; 8,761,478; 8,767,908; 8,779,366; 8,787,620; 8,796,637; 8,809,790; 8,809,793; 8,816,286; 8,818,488; 8,822,933; 8,822,935; 8,828,355; 8,837,799; 8,866,086; 8,884,240; 8,897,518; 8,903,152; 8,907,290; 8,913,810; 8,921,754; 8,921,796; 8,933,409; 8,933,411; 8,934,959; 8,937,285; 8,941,071; 8,942,445; 8,969,815; 8,969,816; 8,969,829; 8,975,907; 8,987,659; 8,992,918; 9,031,300; 9,044,153; 9,063,520; 9,078,622; 9,091,771; 9,140,804; 9,151,851; 9,155,514; 9,176,240; 9,176,241; 9,182,506; 9,207,334; 9,217,795; 9,229,115; 9,244,180; 9,268,033; 9,279,892; 20010001107; 20010040219; 20010056234; 20020113211; 20020145115; 20020195565; 20030038240; 20030047686; 20030047687; 20030057375; 20030062482; 20030105397; 20030116713; 20040016884; 20040026620; 20040084625; 20040129886; 20040164249; 20040183022; 20040195512; 20040200966; 20040210132; 20040258286; 20040260176; 20050004452; 20050015004; 20050023473; 20050031293; 20050035297; 20050061983; 20050082484; 20050082486; 20050113667; 20050117694; 20050129295; 20050151084; 20050156112; 20050167599; 20050230626; 20050242288; 20050253073; 20050253076; 20050285041; 20060029544; 20060108509; 20060116567; 20060138315; 20060145082; 20060163485; 20060197023; 20060202125; 20060231765; 20060237654; 20060261275; 20070010731; 20070040122; 20070057189; 20070090300; 20070116168; 20070131866; 20070181814; 20070205368; 20070221850; 20070263764; 20070267576; 20070269093; 20070270693; 20070278409; 20080011953; 20080069414; 20080118134; 20080128623; 20080135769; 20080137930; 20080156993; 20080164875; 20080197288; 20080203309; 20080210876; 20080219534; 20080230707; 20080237475; 20080240535; 20080253525; 20080253526; 20080253527; 20080253528; 20080253529; 20080253530; 20080253531; 20080253627; 20080260646; 20080284428; 20090018438; 20090072151; 20090072153; 20090074281; 20090110256; 20090146065; 20090159804; 20090161931; 20090161933; 20090169085; 20090175523; 20090220419; 20090224158; 20090236532; 20090257633; 20090262996; 20090264753; 20090302228; 20100010343; 20100012846; 20100014728; 20100033186; 20100046821; 20100072375; 20100074500; 20100076300; 20100078566; 20100078569; 20100084559; 20100098312; 20100104505; 20100108894; 20100108896; 20100108900; 20100116994; 20100135559; 20100140486; 20100148039; 20100152577; 20100166274; 20100182011; 20100189324; 20100198061; 20100200763; 20100219345; 20100219347; 20100220909; 20100230601; 20100246919; 20100252723; 20100268074; 20100294940; 20110001053; 20110018541; 20110073764; 20110079722; 20110105892; 20110116695; 20110117094; 20110133091; 20110142304; 20110142315; 20110142367; 20110150181; 20110174980; 20110210255; 20110212090; 20110215248; 20110218432; 20110220802; 20110228999; 20110248175; 20110248765; 20110253901; 20110272587; 20110278466; 20110291017; 20110299747; 20110301918; 20120018644; 20120019064; 20120022361; 20120022362; 20120022364; 20120061576; 20120068076; 20120068077; 20120093380; 20120112078; 20120114212; 20120129274; 20120138804; 20120148138; 20120155736; 20120157829; 20120157830; 20120193545; 20120223236; 20120241631; 20120265050; 20120271840; 20120290519; 20130009063; 20130009066; 20130020487; 20130028496; 20130032706; 20130032721; 20130032722; 20130131422; 20130131493; 20130136328; 20130149240; 20130193330; 20130240721; 20130256536; 20130256559; 20130284936; 20130310681; 20130315454; 20130320218; 20130327932; 20130334428; 20130341518; 20140003689; 20140021354; 20140021356; 20140029715; 20140048716; 20140062486; 20140064585; 20140079304; 20140110592; 20140175294; 20140183369; 20140194735; 20140200848; 20140206983; 20140224963; 20140257096; 20140275965; 20140276018; 20140276019; 20140276029; 20140316258; 20140330117; 20140334702; 20140336987; 20140367577; 20150001399; 20150001402; 20150001403; 20150021488; 20150036789; 20150057535; 20150065854; 20150076357; 20150090890; 20150117733; 20150119704; 20150160353; 20150177386; 20150192685; 20150199302; 20150212216; 20150219771; 20150262389; 20150285921; 20150285922; 20150289825; 20150307583; 20150323685; 20150331115; 20150355347; 20150370223; 20150374318; 20150378035; 20150380121; each of which is expressly incorporated herein by reference in its entirety.
TABLE 1An incomplete list of devices that represents the current state of technology of human brain PET imaging.DeviceSize/mass/geometry/FOVDetector/photo-SensitivitySpatial resolutionScatterNECEnergyname/yeartransaxial/axialtransducerskcps/kBq/mLTransaxial/axialfraction[kcps] resolutionHitachi2 m × 2 m × 0.5 m/~500CdTe/PMT17.62.3/5.1 mm @ 1 cm23%41 @4.1%kg/ring 310/250 mm4.8/5.9 mm @ 10 cm7.9 kBq/mLStationaryCerePET1 m × 1 m × 0.5 m/~23 kg/ringLYSO/PMTN/A2.0/2.3 mm @ 1 cmN/AN/A13%25/200 mm Portable3.0/? @ 8 cmPET-HATDiameter 1 m × 0.25 m/thick/50GSO/PMT0.724/4.3 @ 1 mm60%0 0.8215%[33]kg/280/44 “Wearable”HRRT dedicated high-resolution 3-[56, 57]dimensional (3D) human braindouble layer39.82.3/3.2 mm45%45PET 312/250LSO/LYSO/PMT2.5/3.4 mm
TABLE 2Properties of selected scintillators suitable for PET.NaI (T1)SrI2:EuLSO:CeLYSO:CeLaBr3:CeMaterial[58][59][60, 61][60, 61][62]Light Output3800030000-120000240002500070000(photons/MeV)Principal Decay 230450-1700404125Time (ns)Wavelength of Max.415435420428360Emission (nm)Attenuation Length3.31.951.21.162.4(511 keV, cm)Density (gm/cc)3.674.557.47.35.0Energy resolution5.6% 2.8% 3% 10% 3%(662 keV)(662 keV)(662 keV)(511 keV)(511 keV)ScintillatorsCe:Gd3Al2Ga3O12Ce:Lu1.8Y0.2SiO5Bi4Ge3O12Ce:LaBr3Lu3Al5O12(Ce:GAGG)(Ce:LYSO)(BGO)LuAGDensity (g/cm3)6.637.17.135.086.73Light Yield57,00034,0008,00075,00022,000(photon/MeV)Decay time (ns)88 (91%); 403003020258 (9%)Peak emission (nm)520420480375310Energy Resolution5.210122.64.2(% @ 662 keV)HygroscopicityNoNoNoYesNoCleavageNoNoNoNoNoMelting point (° C.)1,8502,1501,0507832,043